Multi-path transthoracic defibrillation and cardioversion

ABSTRACT

A defibrillation system for synchronized cardioversion of a patient includes a first housing that includes a measurement circuit configured to receive electrocardiogram (ECG) signals and measure ECG parameters based on the ECG signals, and a first processor configured to analyze the ECG parameters, and initiate communication of a synchronization signal for a second processor for delivery of one or more defibrillation pulses and further includes a second housing that is separate from and external to the first housing and that includes a shock delivery circuit, and the second processor which is configured to receive the communication of the synchronization signal from the first processor, and control the shock delivery circuit to deliver the one or more defibrillation pulses in response to the synchronization signal.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation application under 35 USC § 120 toU.S. patent application Ser. No. 14/930,576, filed on Nov. 2, 2015 whichis a divisional application of U.S. patent application Ser. No.10/960,311, filed on Oct. 6, 2004 and issued as U.S. Pat. No. 9,174,061which is a continuation-in-part application of and claims priority toU.S. application Ser. No. 10/712,308, filed on Nov. 13, 2003 andexpressly abandoned. All subject matter set forth in the abovereferenced applications is hereby incorporated by reference in itsentirety into the present application as if fully set forth herein.

TECHNICAL FIELD

This invention relates to transthoracic defibrillation and cardioversion(i.e., defibrillation or cardioversion performed using electrodesexternal to the thoracic cavity).

BACKGROUND

Normally, electrochemical activity within a human heart causes theorgan's muscle fibers to contract and relax in a synchronized manner.This synchronized action of the heart's musculature results in theeffective pumping of blood from the ventricles to the body's vitalorgans. In the case of ventricular fibrillation (VF), however, abnormalelectrical activity within the heart causes the individual muscle fibersto contract in an unsynchronized and chaotic way. As a result of thisloss of synchronization, the heart loses its ability to effectively pumpblood. Defibrillators produce a large current pulse that disrupts thechaotic electrical activity of the heart associated with ventricularfibrillation and provides the heart's electrochemical system with theopportunity to re-synchronize itself. Once organized electrical activityis restored, synchronized muscle contractions usually follow, leading tothe restoration of effective cardiac pumping.

First described in humans in 1956, transthoracic defibrillation hasbecome the primary therapy for cardiac arrest, ventricular tachycardia(VT), and atrial fibrillation (AF). Monophasic waveforms dominated until1996, when the first biphasic waveform became available for clinicaluse. Attempts have also been made to use multiple electrode systems toimprove defibrillation efficacy. While biphasic waveforms andmultiple-electrode systems have shown improved efficacy relative tomonophasic defibrillation, there is still significant room forimprovement: shock success rate for ventricular fibrillation (VF)remains less than 70% even with the most recent biphasic technology.

Cardiac fibrillation and defibrillation are still poorly understood andseveral hypotheses have been promulgated to explain the mechanisms ofdefibrillation. The concept termed the critical mass hypothesis positsthat a defibrillation shock is successful because it extinguishesactivation fronts within a critical mass of muscle by depolarizing allnon-refractory tissue within a critical mass. The upper limit ofvulnerability (ULV) theory hypothesizes that a shock will be successfulwhen, in addition to terminating ventricular fibrillation (VF)wavefronts by prolonging refractoriness in the myocardium ahead of thewavefront, the shock also must not initiate new fibrillation-causingwavefronts at the border of the shock-depolarized region. A shock may beof sufficient intensity to depolarize the myocardium but not be of highenough intensity to prevent new activation fronts, thus resulting in afailed defibrillation attempt. The critical point hypothesis, related tothe ULV theory, states that a shock must not create a critical pointwhere a critical voltage gradient intersects with a critical point ofrefractoriness. These critical points are the initiation points ofrefibrillation. The “extension of refractoriness” theory states that theshock-induced depolarization of the fibrillating cardiac tissue extendsthe period of refractoriness to incoming VF wavefronts and as a resultterminates VF. Other theories, related to the ULV hypothesis are“progressive depolarization” and “propagated graded (progressive)response cellular depolarization hypothesis”.

The theory of Virtual Electrode Polarization (VEP) describes thephenomena by which, because of current flow within a partiallyconductive medium (the myocardium) contained within another partiallyconductive medium (blood of the cardiac chambers, lungs, interstitialfluids and other organs within the thoracic cavity), myocardialpolarization during defibrillation is characterized by the simultaneouspresence of positive and negative areas of polarization adjacent to eachother. “Phase Singularity” as defined within the context of VEP is acritical point that is surrounded by positively polarized (equivalent to“depolarized” in the conventional electrophysiology nomenclature),non-polarized and negatively polarized (equivalent to “hyperpolarized”)areas. These phase singularities are the source of re-initiation offibrillation. Post shock excitations initiate in the non-polarizedregions between the positively and negatively polarized areas through aprocess termed “break excitation.” The break excitations propagatethrough the shock-induced non-polarized regions termed “excitable gaps”,and if the positively polarized regions have recovered excitability,then a re-entrant circuit at which fibrillation may initiate is formed.The upper limit of vulnerability (ULV) is attained when the areal extentof the excitable gaps is sufficiently minimized, or the shock inducedvoltage gradient is sufficient to cause rapid propagation of theexcitation in the excitable gap, or the extension of refractoriness issufficient to prevent further advance of the break excitations into thedepolarized tissue. With biphasic defibrillation, the second phase ofthe shock tends to nullify the VEP effect by depolarizing the negativelypolarized tissue. Since less energy is needed to depolarize repolarizedtissue than further depolarize already depolarized tissue, effectivebiphasic defibrillation achieves nearly complete depolarization of themyocardium by reversing the negative polarization while maintaining thepositive polarization. There remain, however, excitable gaps withbiphasic waveforms, albeit reduced in scope relative to monophasicwaveforms.

Theoretical approaches to stimulation employing current summation ofmultiple current sources have been used in the past to produce in theoverlap region an additive current or integrated myocardial responsesufficient to cause stimulation or defibrillation while the singularcurrent vectors would not. The approach does not address the issue thatinsufficiently stimulated tissue may remain in the excitable gap thatmay result in refibrillation.

The concept of current equalization has been promulgated as a means ofunderstanding stimulation. The general approach is to equalize thecurrent distribution across the heart and concentrate the current in themuscular areas of the heart. This approach does not address thegeneration of the excitable gap, which will still be present. Asunderstood within the context of the VEP effect, uniform currentdistributions still result in an excitable gap. In fact, a uniformcurrent distribution is not an especially relevant concept within thecontext of a physiological system such as that of the human thorax whereconductances of the organs, muscle, fluids and bone may vary by a factorof 100. Within such a system, current distributions will not be uniform.Even in a simplified, two-conductance system, an applied uniform fieldwill result in a non-uniform current distribution due to the differencein conductances.

The technique of superposed, multiple vector physiologic tissuestimulation has been employed as early as 1948 by Nemec, as disclosed inU.S. Pat. No. 2,622,601, in which a nerve or muscle stimulator isdescribed employing two stimulation waveform generators with multiplesets of electrode. Each waveform is an alternating current electricalsignal with the difference between the two frequencies set to 1-100 Hz.In the areas of tissue that are exposed to currents from bothsources—the regions of current superposition—a beat frequency equal tothe frequency difference will be generated that is capable ofstimulating the physiological tissue. U.S. Pat. No. 3,774,620 added theconcept of superposition of two or more AC currents that by themselveshave no stimulative effect, the currents differing from each other by alow value, with an optimum interference in the treatment area. Similarmethods were employed in U.S. Pat. Nos. 3,774,620, 3,895,639, 4,023,574,and 4,440,121. In these and much of the subsequent art, the regions ofinterest were those areas where the current from the multiple sourcesoverlapped. The summation current in the overlap region would result ina beat frequency or additive current, which would be sufficient to causestimulation while the singular current vectors would not.

The earliest cardioverters and defibrillators generated either a singleburst of alternating current or a single pulse for application to theheart to cause cardioversion or defibrillation. However, the use ofmultiple pulses to accomplish cardioversion or defibrillation has alsobeen in extensively researched. U.S. Pat. No. 3,605,754 discloses anearly double pulse heart defibrillator employing two capacitors that aresuccessively discharged between a single pair of electrodes.Multiple-electrode systems have been employed for implantable pacemakersand defibrillators. For example, sequential pulse multiple electrodesystems are disclosed in U.S. Pat. Nos. 4,291,699, 4,641,656, 4,708,145,4,727,877 4,932,407, and 5,107,834. Sequential-pulse systems operatebased on the assumption that sequential defibrillation pulses, deliveredbetween differing electrode pairs have an integrative effect, due to thenon-linear action potential response of cardiac tissue, such that theoverall energy requirements to achieve defibrillation are less than theenergy levels required to accomplish defibrillation using a single pairof electrodes. An alternative approach to multiple-electrode,sequential-pulse defibrillation is disclosed in U.S. Pat. No. 4,641,656.One electrode pair may include a right ventricular electrode and acoronary sinus electrode, and the second electrode pair may include aright ventricular electrode and a subcutaneous patch electrode, with theright ventricular electrode serving as a common electrode to bothelectrode pairs. An alternative multiple-electrode, simultaneous-pulsesystem is disclosed in U.S. Pat. No. 4,953,551, employing rightventricular, superior vena cava and subcutaneous patch electrodes. U.S.Pat. No. 4,953,551 discloses simultaneous delivery of pulses between thesuperior vena cava and the right ventricle and between the rightventricle and a subcutaneous electrode. In U.S. Pat. No. 5,163,427, twocapacitor banks are provided which are simultaneously charged and thensuccessfully or simultaneously discharged between different pairs ofelectrodes.

French Patent No. 2,257,312 discloses sequential pulse defibrillatorsemploying multiple electrodes arranged in and around the heart. In thatdisclosure, alternating current (AC) defibrillation pulses aresequentially delivered such that each successively activated electrodepair defines a pulse vector, and such that the pulse vectors scan in arotational fashion through the heart tissue. Pulses are deliveredimmediately following one another, or may overlap one another for someunspecified period. U.S. Pat. No. 5,324,309 describes overlapping dualpathway pulses where there is an intermediate current vector during theoverlap period. U.S. Pat. No. 5,766,226 describes a similarconfiguration in which the intermediate current vector changes directionand is made to cycle back and forth during the shock pulse. A similarconfiguration is described in U.S. Pat. No. 5,800,465. U.S. Pat. No.5,330,506 describes a multi-pathway pacing method where each individualpath is a subthreshold stimulus while the current level in the region ofsuperposition is suprathreshold. The current vector in the region ofsuperposition can be steered by varying the timing of the individualpulse onsets. U.S. Pat. No. 5,431,688 describes a multi-electrode,focused waveform, with interposed pulse trains. These techniques havesimilar deficits in that, while they are able to reduce the excitablegap to some extent via the rotating vector produced by the overlappingof the currents, regions of excitable gap will remain that can stilltrigger refibrillation.

U.S. Pat. No. 6,148,233 describes a multi-contact electrode composed ofmultiple small active areas, each active area of a size too small todefibrillate. Each active area is connected to the same current source.Division of the electrode into plural active areas is intended toprovide a means of reducing skin sensitization from long-term wear ofthe electrodes.

SUMMARY

In a first aspect, the invention features a transthoracic defibrillatorfor external defibrillation, the defibrillator comprising three or moreelectrodes configured to be attached to the thorax of a patient toestablish at least two electrical paths across the thoracic cavity andthrough the heart of the patient, cables to connect the three or moreelectrodes to a defibrillator circuit contained in a defibrillatorhousing, wherein the defibrillator circuit has the capability to delivera different defibrillation waveform across each of the at least twoelectrical paths.

Preferred implementations of this aspect of the invention mayincorporate one or more of the following. The different defibrillationwaveforms may differ in at least one waveform parameter. Thedefibrillator circuit may have the capability to deliver the samedefibrillation waveform across each of the at least two electrical,paths. The defibrillator circuit may include a processing unit fordetermining an transthoracic impedance distribution and for selectingthe waveform parameter of the at least two electrical paths based on thetransthoracic impedance distribution. The transthoracic impedancedistribution may be two dimensional. The transthoracic impedancedistribution may be determined by measuring impedances between locationson the thorax. Measuring impedances between locations on the thorax maycomprise measuring impedances between the electrodes. The transthoracicimpedance distribution may be measured using electrical impedancetomography (EIT). The transthoracic impedance distribution may bemeasured using an imaging technique to determine positions of tissueregions, and computing the transthoracic impedance distribution from thepositions of tissue regions and resistivities of the tissues. Theimaging technique may comprise ultrasound imaging. The imaging techniquemay employ at least one transducer element integrated into adefibrillation pad supporting at least one of the electrodes. At leastone parameter of each waveform may be one of tilt, duration, current, orvoltage. The waveforms may be biphasic. The waveforms may be monophasic.The waveforms may be multiphasic. The waveforms may be interlaced. Atleast one parameter of each waveform may be one of tilt, duration,current, voltage, first phase duration, second phase duration, firstphase average current. The waveforms across different electrical pathsmay be overlapping in time by at least 1 millisecond but by less than 80percent of the duration of the shortest of the waveforms. The waveformsacross different electrical paths may be delivered simultaneously. Thewaveforms across different electrical paths may be deliveredsequentially without overlapping in time. At least one waveformparameter of each waveform may be adjusted to achieve substantially thesame defibrillation efficacy for each electrical path. At least onewaveform parameter of each waveform may be adjusted to achieve aselected current density distribution at the heart. At least onewaveform parameter of each waveform may be adjusted to make the currentdensity distribution at the heart more uniform than would be the case ifthe waveform parameter were the same for each of the electrical paths.The current density may be either peak or average current density. Atleast two electrodes positioned on the same side of the thorax may becombined into a unitary electrode pad that is adhered to and removedfrom the patient as one unit. There may be at least four electrodes, twoon each side of the thorax, and two electrodes on each side of thethorax may each be combined into a unitary electrode pad that is adheredto and removed from the patient as one unit. The area of each of theelectrodes through which the waveforms are delivered may be less than 70percent of the projected area of the heart, and the sum of the areas ofthe electrodes on the same side of the thoracic cavity may be greaterthan 80 percent of the projected area of the heart. The determination ofa transthoracic impedance distribution may occur at the time of or justprior to delivery of the defibrillation waveforms.

In a second aspect, the invention features a method of externalelectromagnetic stimulation of the interior of the body, the methodcomprising applying three or more electrodes to the exterior of thepatient to establish at least two electrical paths across the interiorof the patient, determining impedance information representative of animpedance distribution across the interior of the body, delivering anelectromagnetic waveform across each of the at least two electricalpaths, wherein at least one parameter of the waveform is selected usingthe impedance information to produce a selected current densitydistribution at one or more locations within the interior of the body.

Preferred implementations of this aspect of the invention mayincorporate one or more of the following. The electromagneticstimulation may be for defibrillation or cardioversion of the heart, theimpedance distribution may be across the thorax, and the current densitydistribution may be at the heart. Determining impedance information maycomprise electrical impedance tomography (EIT). Determining theimpedance information may comprise imaging the body to determinepositions of tissue regions, and computing the transthoracic impedancedistribution from the positions of tissue regions and resistivities ofthe tissues. Imaging may comprise ultrasound imaging. The selectedcurrent density may be configured to deliver additional current densityto selected myocardial regions expected to receive insufficientmyocardial stimulation from a defibrillation shock. The selectedmyocardial regions may include an excitable gap region. The currentdensity distribution may be selected using a model (e.g., a bidomainmodel) that predicts the regions expected to receive insufficientmyocardial stimulation. The electromagnetic waveforms providing theselected current distribution may be delivered following adefibrillation pulse expected to deliver insufficient myocardialstimulation to the selected myocardial regions. ECG signals may bedetected from a plurality of electrodes, and epicardial voltages may beestimated from the detected ECG signals. The one or more locations atwhich the selected current density distribution is produced may be areasof the heart, and the epicardial voltages may be used to determine theareas of the heart. Morphologies may be selected for the electromagneticwaveforms to produce the selected current density distribution at theareas of the heart. The areas of the heart may be areas in front of anadvancing activation wavefront, and the selected current densitydistribution may be sufficient to quench the wavefront.

In a third aspect, the invention features a method of performingtransthoracic defibrillation, comprising attaching three or moreelectrodes to the thorax of a patient to establish a plurality ofelectrical paths across the thoracic cavity and through the heart of thepatient, and delivering a defibrillation waveform across each of atleast two of the electrical paths, wherein the area of each of theelectrodes through which the waveforms are delivered is less than 70percent of the projected area of the heart, and the sum of the areas ofthe electrodes on the same side of the thoracic cavity is greater than80 percent of the projected area of the heart.

Preferred implementations of this aspect of the invention mayincorporate one or more of the following. The invention furthercomprises measuring an electrical, electrocardiographic, physiological,or anatomical parameter of the patient, and delivering defibrillationwaveforms that under at least some circumstances may be different fordifferent electrical paths, with at least one parameter of each waveformbeing dependent on the measured parameter. Measuring may comprisedetermining a transthoracic impedance distribution. The area of each ofthe electrodes through which the waveforms are delivered may be lessthan 60 percent of the projected area of the heart. The area of each ofthe electrodes through which the waveforms are delivered may be lessthan 50 percent of the projected area of the heart. The sum of the areasof the electrodes on the same side of the thoracic cavity may be greaterthan 90 percent of the projected area of the heart. The sum of the areasof the electrodes on the same side of the thoracic cavity may be greaterthan 100 percent of the projected area of the heart. The electrodes maybe positioned in anterior and posterior locations, so that theelectrical paths extend between the anterior and the posterior of thepatient's thorax. The electrodes may be positioned at lateral locations,so that the electrical paths extend between left and right sides of thepatient's thorax. The waveforms may be multiphasic. The waveforms may bemonophasic. At least two of the electrodes may be combined in oneunitary electrode pad that is applied and removed from a patient as aunit. There may be a seam line between areas of the pad in whichelectrodes are supported, with the seam line being constructed so thatthe pad can be folded without creasing the areas in which electrodes aresupported. There may be a multiplicity of electrodes arranged on theunitary electrode pad. The multiplicity of electrodes may be arranged toincrease packing density. The electrodes may be arranged in the form ofa polygon tessellation. The tessellation may be a regular tessellationcomprising regular polyhedra symmetrically tiling a plane. The polyhedramay be one of a triangle, square, or hexagon.

In a fourth aspect, the invention features a method of performingtransthoracic defibrillation, comprising attaching three or moreelectrodes to the thorax of a patient to establish at least twoelectrical paths across the thoracic cavity and through the heart of thepatient, delivering a biphasic or multiphasic defibrillation waveformacross each of the at least two electrical paths, wherein under at leastsome circumstances the multiphasic waveforms delivered are different forthe at least two electrical paths.

Preferred implementations of this aspect of the invention mayincorporate one or more of the following. The invention furthercomprises determining a transthoracic impedance distribution across theat least two electrical paths, and delivering the biphasic ormultiphasic waveforms with at least one parameter of each multiphasicwaveform being dependent on the transthoracic impedance distribution.There may be two pairs of electrodes, with one electrode of each pairlocated on generally opposite surfaces of the thorax. One electrode ofeach pair may be located on the anterior and the posterior surfaces ofthe thorax. The invention further comprises a pair of bridge circuits,one bridge circuit for generating each of the biphasic or multiphasicwaveforms. The biphasic or multiphasic waveforms may be delivered so asto overlap in time. The biphasic or multiphasic waveforms may besimultaneous. The biphasic or multiphasic waveforms may be sequential.

In a fifth aspect, the invention features a defibrillation electrodecomprising a first electrical wire for conveying a defibrillation pulseto or from the electrode, a metallic layer connected to the electricalcable, a conductive, skin-contacting layer for conveying the pulse fromthe metallic layer to the skin, an ultrasound sensor, and a secondelectrical wire for connecting the ultrasound sensor to an ultrasoundimaging circuit.

In a sixth aspect, the invention features a method of performingtransthoracic defibrillation, comprising attaching three or moreelectrodes to the thorax of a patient to establish a plurality ofelectrical paths across the thoracic cavity and through the heart of thepatient, using at least two different defibrillation circuits togenerate at least two generally different defibrillation waveforms,delivering one of the at least two different defibrillation waveformsacross each of the at least two electrical paths, and synchronizingdelivery of the at least two defibrillation waveforms by communicationsbetween the at least two different defibrillation circuits.

Preferred implementations of this aspect of the invention mayincorporate one or more of the following. The defibrillation circuitsmay each comprise a processor, an energy delivery circuit, and aswitching circuit. The defibrillation circuits may be contained inseparate housings, and the communications occurs between the housings.The switching circuit may be capable of generating a biphasic ormultiphasic defibrillation waveform. The synchronizing delivering mayinclude analog communication between the defibrillation circuits. Thesynchronizing delivering may include digital communication between thedefibrillation circuits. The defibrillation circuits may be contained inseparate housings. The defibrillation waveforms may deliver primarilyelectrical current. The defibrillation waveforms may deliver primarily amagnetic field. There may be an energy delivery circuit comprising oneor more capacitors, a charging circuit for charging the one or morecapacitors, and a switching circuit coupled to the one or morecapacitors. An additional switch may be provided for decoupling thecapacitor from the charging circuit prior to delivery of the waveform.The switching circuit may be configured as a Class D amplifier. Theswitching circuit may be configured as a Class B amplifier. Theswitching circuit may be configured as a Class AB amplifier. Theinvention further comprises delivering diaphragmatic stimulation. Atleast one diaphragmatic electrode may be provided for delivering thediaphragmatic stimulation. At least two of the defibrillation electrodesand at least one diaphragmatic electrode may be combined in one unitaryelectrode pad that is applied and removed from a patient as a unit. Adevice for delivering chest compressions may be provided. The device fordelivering chest compressions may comprise a compression bandsurrounding the thorax. The device for delivering chest compressions maycomprise a piston-driven device. A physiological parameter may bemeasured, and a prediction of defibrillation success based on analysisof the measured physiological parameter, and a coordinated delivery ofdefibrillation and chest compressions may be provided based on theprediction. The coordinated delivery of defibrillation and chestcompressions may be manual, advisory, semi-automated, or fullyautomated. Diaphragmatic stimulation for assisted breathing may also beprovided. Cardiac pacing may also be provided. Delivery of a seconddefibrillation waveform may be initiated after a delay intervalfollowing initiation of delivery of a first waveform. It has been foundthat delays in the range of 15-40 milliseconds result in an increase inthe probability of refibrillation. The delay interval may be less than15 milliseconds. Alternatively, the delay interval may be greater than40 milliseconds but less than 200 milliseconds.

In a seventh aspect, the invention features a transthoracicdefibrillator that comprises three or more electrodes configured to beattached to the thorax of a patient to establish at least two electricalpaths across the thoracic cavity and through the heart of the patient;cables to connect at least some of the electrodes to a defibrillatorcircuit contained in a defibrillator housing, wherein the defibrillatorcircuit has the capability to deliver a first defibrillation waveformacross a first electrical path and a second defibrillation waveformacross a second electrical path, and wherein the locations andconfigurations of the electrodes and the first and second waveforms areconfigured so that the first waveform has a first current vector at theheart that is substantially aligned with the long axis of the fibers ofa first portion of the heart and so that the second waveform has asecond current vector at the heart that is substantially aligned withthe long axis of the fibers of a second portion of the heart.

Preferred implementations of this aspect of the invention mayincorporate one or more of the following. The defibrillator circuit mayhave the capability to make the first and second defibrillation waveformdifferent from one another. A three-dimensional imaging method may beused to determine the orientation of the myocardial fibers. The imagingmethod (e.g., MRI) may include the capability of measuring currentdensities along with three dimensional volume images. Current may beinjected during an MRI image generation by an external source to provideimages of current flow in myocardial fibers. The defibrillator may beimplanted, a first and second of the three or more electrodes may beprovided by two separate housings each containing at least somedefibrillator circuitry, the housings being electrically connected toone another, and a third and fourth of the three or more electrodes maybe positioned outside the thoracic cage and connected electrically toone of the two separate housings. All of the electrodes used fordelivering a defibrillation shock may be located outside the thoraciccavity. There may be a common electrical cable between the two separatehousings, the common electrical cable comprising at least one conductorfor carrying stimulation pulses to one of the electrodes and at leastone conductor for carrying communications between the separate housings.The third electrode may comprise at least two intercostal electrodeselectrically connected in common, with each intercostal electrodepositioned in the vicinity of an intercostal space.

Among the many advantages of the invention (some of which may beachieved only in some of its various aspects and implementations) arethat the invention can provide a transthoraciccardioversion/defibrillation system that results in reduced areal extentand effects of the excitable gap during defibrillation and providesimproved efficacies relative to prior art. In some implementations, theexcitable gap areal extant is reduced by increasing the combined area ofthe electrodes on each side of the thorax (e.g., making the combinedarea approximately equal to or larger than the area of the heartprojected onto those electrodes). Because the transthoracic impedancevaries considerably across the surface of the chest, affected by thebone, skeletal muscle and cartilage of the sternum and ribs and thedisparate conductances of the myocardium, blood and lung, someimplementations of the invention are capable of adjusting at least oneof the parameters of the waveforms such as duration, waveform shape oramplitude based on a determination of the transthoracic impedancedistribution. Such a determination might be as simple as an impedancemeasurement between the electrodes of the electrode pair, but might alsoinclude electrical impedance tomography, ultrasonic imaging or otherimaging method to determine more accurate locations of the heart, lungsand skeleton. In some implementations, the variation in transthoracicimpedance may be addressed by configuring the power sources deliveringwaveforms as independent current sources with the current set to adesired value across a range of physiological impedances.

Conductances at the body surface do not vary nearly as much as those ofthe internal organs, fluids, muscle and bone. Thus, the typical priorart single-point impedance measurement (e.g., at the body surface usingthe same electrodes that deliver the defibrillation current) is unableto estimate the impedance distribution within the thoracic cavity.Waveform shape, amplitude, or duration has been varied depending on suchsingle-point impedance measurements, and this approach can have animpact on the efficacy of the defibrillation pulse, but its effect islimited by the fact that it does nothing to alter the currentdistribution in and around the heart. In preferred implementations, theinvention determines the impedance (resistivity) distribution of thethorax in at least two dimensions, and uses the impedance distributionto determine the waveform parameters for each electrical path (currentvector). E.g., the amplitude of the defibrillation pulse for eachelectrode pair can be independently adjusted to achieve a desiredcurrent distribution in and around the myocardium. Using suchimplementations of the invention, the current actually delivered to theorgans themselves can be controlled at the surface of the body on asfine a level of detail as determined by the number, location and size ofthe electrodes located on the body surface.

Depending on the particular implementation, the invention is capable ofimproving the performance of any of the known types of defibrillationwaveforms: monophasic, biphasic, or multiphasic. Preferably, thewaveforms of the individual vectors are synchronized, but the inventionis also capable of improving the performance of sequentially pulsed,multi-electrode defibrillation systems.

Some implementations of the invention measure the electrical,electrocardiographic, physiological, or anatomical parameters of thepatient along an axis substantially similar to the axis of an electrodepair at the time of or just prior to defibrillation, and use themeasurements to control the waveform parameters to improve efficacy.

Some implementations of the invention have current pathways that areindependently measurable and controllable, but in other, simplerimplementations, a waveform parameter of at least one of the currentpathways is controlled as a function of the electrical,electrocardiographic, physiological or anatomical parameters of thepatient.

Some implementations of the invention provide for synchronizing in amaster/slave fashion multiple defibrillators that individually canfunction as standard monophasic or biphasic defibrillators.

The multiple electrodes may be implemented as active electrode areasintegrated into a single pad, with one pad applied to each side of thethorax, thereby achieving a two-pad, easy-to-use system. The integratedpads may include anatomical markings such as correctly placed drawingsof the sternum, sternal notch, or nipples to provide the clinician withthe ability to more accurately place the electrodes on the patient. Theelectrodes may include sensors such as ultrasonic imaging, impedance,pulse oximetry, end-tidal carbon dioxide, blood pressure, velocitysensing, acceleration sensors.

The active electrode areas of the integrated pad may be configured toprovide optimum spacing by placing them in a close-packed hexagonal,rectangular, or concentric configuration or other tessellations.

Other features and advantages of the invention will be apparent from thedrawings, detailed description, and claims.

DESCRIPTION OF DRAWINGS

FIG. 1 is a schematic of the circuitry for a biphasic defibrillatorimplementation.

FIG. 2 is a block diagram of the implementation of FIG. 1.

FIG. 3 is a plot of a biphasic waveform produced by the implementationof FIG. 1.

FIG. 4 is a cross section of the thorax at the elevation of the heartused in finite element modeling showing the finite element meshdecomposition.

FIG. 5a is a simplified version of FIG. 4 with electrode locations andcoordinate axes and plans.

FIG. 5b is a simplified version of FIG. 4 showing a preferred set ofelectrode locations.

FIG. 5c is a simplified version of FIG. 4 showing an alternative set ofelectrode locations.

FIG. 5d is a simplified version of FIG. 4 showing a further alternativeset of electrode locations.

FIG. 6 shows an isoconductance plot of the anterior thorax.

FIG. 7 shows the placement of the anterior electrodes of FIG. 5brelative to the heart.

FIGS. 8a and 8b show the placement of the electrodes of FIG. 5 b,

FIG. 9 shows an example of an annular electrode configuration.

FIG. 10 shows examples of electrodes arranged in regular tessellations.

FIG. 11 shows examples of electrodes arranged in semi-regulartessellations.

FIG. 12 shows examples of electrodes arranged in demi-regulartessellations.

FIG. 13 shows a simulation of the effects on the heart due to amonophasic defibrillation pulse as modeled using the Virtual ElectrodeTheory.

FIG. 14a, 14b depict, diagrammatically, what occurs when the area of theelectrodes is varied.

FIG. 15 is an electrical schematic of an H-Bridge Class D configurationcircuit.

FIG. 16 is an electrical schematic of circuitry driving an H-bridgeClass D configuration circuit.

FIG. 17 are waveforms produced by the H-bridge Class D configurationcircuit.

FIG. 18 is a plan view of an integrated defibrillation pad including anultrasonic gel window for application of an ultrasonic probe.

FIG. 19 is a block diagram of an implementation with dualdefibrillators.

FIG. 20 is a plan view of two integrated defibrillation electrode pads,electrically connected with a common connector in a dual defibrillatorsystem.

FIG. 21 is a perspective view of a patient to which an integrateddefibrillation pad with anatomical markings has been applied.

FIG. 22 is a block diagram of an integrated resuscitation systemimplementation.

FIG. 23 is a decision flowchart for the integrated resuscitation system.

FIGS. 24A and 25A show two different examples of the spatialdistribution of epicardial voltages across a region of the heartimmediately subsequent to defibrillation shock. Different shades of grayrepresent different voltage levels, with the darkest region at −90 mV(−50 mV in FIG. 25A) at the extreme left in the figure indicating theregion of greatest depolarization, and the darkest region at 0 mV (+10mV in FIG. 25A) at the extreme right indicating the region of greatesthyperpolarization. The white region and the bands immediately adjacentare the excitable gap. Isopotential contour lines are drawn 5 mV aparton the basis of the assumption that the normal resting potential is 285mV and action potential amplitude is 100 mV.

FIGS. 24B and 25B show two different examples of the spatialdistribution of epicardial phase across a region of the heartimmediately following a defibrillation shock. Different shades of grayrepresent different isochronous regions (i.e., regions at approximatelythe same cardiac phase). The regions are 10 msec apart, with the regionin white representing the time immediately after the defibrillationshock (0 msec), and the region in the darkest gray representing 120 msecafter the shock.

FIGS. 24C and 25C are ECG signals representing possible behaviorimmediately after a defibrillation pulse, and show onset of tachycardia.

FIG. 25D is the same as FIG. 25A but with the excitable gap region shownas a cross-hatched area.

FIG. 25E is the same as FIG. 25B but with a cross hatched arearepresenting the area of an applied quenching stimulus.

FIG. 26 provides a simplified, perspective view of the ventricles,showing the generally spiral orientation of the muscle fibers.

FIG. 27 provides a simplified, perspective view of the ventricles,showing two possible current vectors (for a two-vector approach)generally aligned with the muscle fibers.

FIGS. 28A and 28B show an implanted defibrillator and electrodes thatprovide another implementation of the two vector approach in which thevectors are aligned with muscle fibers.

FIG. 29 shows an electrical schematic of the implementation of FIG. 28B.

DETAILED DESCRIPTION

There are a great many possible implementations of the invention, toomany to describe herein. Some possible implementations that arepresently preferred are described below. It cannot be emphasized toostrongly, however, that these are descriptions of implementations of theinvention, and not descriptions of the invention, which is not limitedto the detailed implementations described in this section but isdescribed in broader terms in the claims.

One implementation of the invention is depicted in FIG. 1. Thedefibrillation waveform delivered to the patient is a biphasic ormultiphasic waveform as described in U.S. Pat. No. 6,096,063. Asdescribed in that patent, the electromagnetic (EM) energy delivery means1 is comprised of storage capacitors 2, 3 which are charged to atherapeutically effective voltage by a charging circuit 4 under controlof the processing means 5 while relays 6, 7, 8 and 9 and the H-Bridges10, 11 are open. As a means of reducing both size and cost, chargingcircuit 4 is used to charge both storage capacitors 2, 3 simultaneously.The first electrode pair 1 and the second electrode pair 2 are selfadhesive pads, such as STAT-PADZ (ZOLL Medical Chelmsford Mass.), thatare adhered to the patient's chest 3, shown in cross-section in FIG. 1.

Upon determination by processing means 5, using any existing methodsknown to those skilled in the art, of the appropriate time to deliverthe defibrillation energy to the patient, relay switches 12, 13, 14 and15 are opened, and relay switches 6, 7, 8 and 9 are closed. Then, theelectronic switches 16, 17, 18, and 19 of H-bridge 10 and 24, 25, 26,and 27 of H-bridge 11 are closed to allow electric current to passthrough the patient's body in one direction, after which electronicswitches 16, 17, 18, and 19 of H-bridge 10 and 24, 25, 26, and 27 ofH-bridge 11 are opened and 20, 21, 22, and 23 of H-bridge 10 and 28, 29,30 and 31 of H-bridge 11 are closed to allow the electric current topass through the patient's body in the other direction. Relay switches12, 13, 14 and 15 are combined in double-pole double-throw configuration(DPDT) to reduce size and cost. DPDT relay 12, 13 serves the purpose ofisolating the current sources for the electrode pairs during discharge.Electronic switches 16-31 are controlled by signals from respectiveopto-isolators, which are, in turn, controlled by signals from theprocessing means 5. As shown in FIG. 2, processing means 5 is preferablya microprocessor, such as a Hitachi SH-3 40 combined with a read onlymemory device (ROM) 41, random access memory (RAM) 42, Clock 43, realtime clock 44, analog-to-digital 45 and digital-to-analog 46 converters,power supply 47, reset circuit 48, general purpose input/output 49, anduser interface in the form of a display 49 and input keys 50 and othercircuitry known to those skilled in the art. A measurement means 52 isprovided for measurement of electrical, electrocardiographic,physiological or anatomical parameters of the patient, the processingmeans 5 controlling the waveform parameters of at least one of thedischarge pathways based on this measurement. Relay switches 6, 7, 8,and 9 which are also controlled by the processing means 5, isolatepatient 3 from leakage currents of H-bridge switches 16-31 which may beabout 500 microamperes.

Resistive circuits 55, 56 that include series-connected resistors 57,58, 59 and 60, 61, 62, respectively, are provided in the current path,each of the resistors being connected in parallel with shorting switch63-68 controlled by processing means 5. The resistors are preferably ofunequal value and stepped in a binary sequence such that with thevarious combinations of series resistance values, there are 2^(n)different combinations, where n is the number of resistors. Immediatelyprior to delivering the therapeutic defibrillation energy a smalleramplitude “sensing” pulse is delivered by closing H-bridge switches16-19 and 24-27 and the resistor shorting switches 63-68 are all open sothat current passes through the resistors in series. The current sensingtransformers 69 and 70 sense the current that passes through the patientthrough their respective electrode pairs 1 a, 1 b, 2 a and 2 b, fromwhich the processing means 5 determines the resistance of the patient 3.

The initial sensing pulse is integral with, i.e., immediately followedby, a biphasic defibrillation waveform, and no re-charging of storagecapacitor occurs between the initial sensing pulse and the biphasicdefibrillation waveform. If the patient resistance sensed during theinitial sensing pulse is low, all of the resistor-shorting switches63-68 are left open at the end of the sensing pulse so that all of theresistors 57-62 remain in the current path (the resistors are thensuccessively shorted out during the positive phase of the biphasicdefibrillation waveform in the manner described below in order toapproximate a rectilinear positive phase). Thus, the current at thebeginning of the positive first phase of the biphasic defibrillationwaveform is the same as the current during sensing pulse. If the patientresistance sensed during the sensing pulse is high, some or all of theresistor-shorting switches 63-68 are closed at the end of the sensingPulse, thereby shorting out some or all of the resistors.

Thus, immediately after the sensing pulse, the biphasic defibrillationwaveform has an initial discharge current that is controlled bymicroprocessor 46, based on the patient impedance sensed bycurrent-sensing transformer 69, 70. The current level of the sensingpulse is always at least 50 percent of the current level at thebeginning of positive first phase, and the sensing pulse, like thedefibrillation pulse, is of course a direct-current pulse.

By appropriately selecting the number of resistors that remain in thecurrent path, the processing means reduces (but does not eliminate) thedependence of peak discharge current on patient impedance, for a givenamount of charge stored by the charge storage device. For a patientimpedance of 15 ohms, the peak current is about 25 amperes, whereas fora patient impedance of 125 ohms, the peak current is about 12.5 amperes(a typical patient is about 75 ohms.)

During the positive phase of the biphasic waveform, some or all of theresistors 57-62 that remain in series with the patient 3 aresuccessively shorted out. Every time one of the resistors is shortedout, an upward jump in current occurs in the waveform, thereby resultingin the sawtooth ripple shown in the waveform of FIG. 3. The ripple tendsto be greatest at the end of the rectilinear phase because the timeconstant of decay (RC) is shorter at the end of the phase than at thebeginning of the phase. Of course, if all of the resistors have alreadybeen shorted out immediately after the end of the sensing pulse, thepositive phase of the biphasic waveform simply decays exponentiallyuntil the waveform switches to the negative phase.

As is shown in FIG. 3, at the end of the positive phase, the currentwaveform decreases through a series of rapid steps from the end of thepositive phase to the beginning of negative phase, one of the stepsbeing at the zero crossing. Processing means 5 accomplishes this by (1)successively increasing the resistance of resistive circuit 55, 56 infixed increments through manipulation of resistor-shorting switches57-62, then (2) opening all of the switches in H-bridges 10-11 to bringthe current waveform down to the zero crossing, then (3) reversing thepolarity of the current waveform by closing the H-bridge switches thathad previously been open in the positive phase of the current waveform,and then (4) successively decreasing the resistance of resistancecircuit 55, 56 in fixed increments through manipulation ofresistor-shorting switches 57-62 until the resistance of resistancecircuit 55, 56 is the same as it at the end of the positive phase.

In one implementation a variable resistor 71, 72 is provided in serieswith the other resistors 57-62 to reduce the sawtooth ripple. Every timeone of the fixed-value resistors 57-62 is shorted out, the resistance ofvariable resistors 71, 72 automatically jumps to a high value and thendecreases until the next fixed-value resistor is shorted out. Thistends, to some extent, to smooth out the height of the sawtooth ripplefrom about 3 amps to about 0.1 to 0.2 amps, and reduces the need forsmaller increments of the fixed-value (i.e., it reduces the need foradditional fixed-value resistor stages).

A cross-sectional view of the human thorax is shown in FIG. 4. Each ofthe constituent tissues are subdivided into cells for use in finiteelement simulations of the fields and currents generated bydefibrillation pulses. Electrode pairs 1 a, 1 b, 2 a, and 2 b are alsodepicted in the figure. FIG. 5a-d depicts a simplified version of thecross section of FIG. 4. A line is defined in the figure, the CardiacCenter of Mass (CCOM) line 75, which runs through the CCOM point 76 andis parallel to the patient's back. In preferred implementations, atleast one, and preferably two, electrodes are located posterior to theCCOM line. Additionally, the midpoint/COM (MCOM) line 77 is the linedefined by midpoint of the lateral extent (MLE) of the posteriorelectrode or electrodes 78, 79 and the CCOM point 76. The electrodeplane 81, is defined by the plane resulting in the least mean squarederror distance to the centroids 82 of the electrodes distal 84 to theMLE 78. There is further defined a Projected Cardiac Area (PCA), that isthe area in the electrode plane 81 of the shape formed by theintersection of the electrode plane 81 with the locus of lines 80parallel to the MCOM line 77 and tangent to the surface of the heart 83.The area, shape and position of the electrodes are such that the area ofeach individual electrode is less than 70% of the PCA and the sum of theareas of the electrodes distal 84 to the MLE 78 is greater than 80%, andpreferably 100%, of the PCA.

In a preferred implementation, the electrodes are positioned as shown inFIGS. 5b , 7, 8 a and 8 b. FIG. 7 shows the relative location of theelectrodes 1 a, 2 a and the thoracic cage and the heart 4. FIGS. 8a and8b show the electrode placement on a typical patient. FIG. 5c depicts alateral placement of the electrode pairs. In another implementation, theelectrodes may be so configured as concentric, as shown in FIG. 9. Theelectrodes may also be placed so that the current pathways areessentially parallel, as shown in FIG. 5d (in which the locations ofelectrodes 1 b and 2 b have been reversed from FIG. 5b ).

The conductances of the various tissues as shown in FIG. 4 areapproximately as follows:

Tissue type Conductivity (ohms-cm) Skin 3.4 Blood 6.5 Lung 0.7 SkeletalMuscle 1.5 (transverse) 4.2 (longitudinal) Fat 0.5 Cardiac Muscle 7.6Bone 0.06

Conductivities of the various tissues can vary by as much as a factor of100. To accommodate this, waveform parameters of the energy delivered toeach of the discharge pathways is independently controllable. Forexample, this may be accomplished in the just-described implementationby providing two high voltage capacitors 2, 3 and by appropriatelyswitching the resistors 57-62 that remain in series with the patient 3.By appropriately selecting the number of resistors that remain in thecurrent path, the dependence of peak discharge current on patientimpedance can be reduced (but not eliminated), for a given amount ofcharge stored by the charge storage device. For example, for a patientimpedance of 15 ohms, the peak current is about 25 amperes, whereas fora patient impedance of 125 ohms, the peak current is about 12.5 amperes(a typical patient is about 75 ohms.)

Alternatively, independent control may also be achieved by providingonly one high voltage capacitor for more than one of the electrode pairswhile still providing separate resistor networks 57-59 and 60-62 foreach current pathway. Another waveform parameter that may be adjusted iswaveform duration, which is controllable by switch networks 10, 11. Theaverage first phase current can also be independently adjusted, e.g., byproviding a second charging circuit 4 to charge a second group of one ormore capacitors to a voltage independent from the first group of one ormore capacitors. Waveform parameters for independent adjustment include,but are not limited to, tilt, duration, first phase duration, secondphase duration, current, voltage, and first phase average current.

As can be seen in the isoadmittance curves shown in FIG. 6, theconductances of the internal organs, muscle and bone vary significantly,much more so than do conductances at the body surface. In a preferredimplementation, electrical impedance tomography (EIT) is used todetermine these internal conductances or impedances. Electricalimpedance tomography (EIT) is used to determine the resistivitydistribution of the thorax in at least two dimensions, and thecalculated resistivity distribution is then used to determine thewaveform parameters for each current vector. For example, the amplitudeof the defibrillation pulse for each electrode pair can be independentlyadjusted to achieve the optimal current distribution in and around themyocardium. Using such a method, the current actually delivered to theorgans themselves can be controlled at the surface of the body on asfine a level of detail as determined by the number, location and size ofthe electrodes located on the body surface.

In the most basic implementation, only three electrodes with threepossible electrode pairs is sufficient to use EIT methods to determinewaveform parameters. In the preferred implementation shown in FIG. 5b ,four electrodes are used, for a total of six [(n−1)!] possible electrodepairs. This number is chosen for ease of implementation and cost;implementations with more electrode pairs are possible.

The EIT system is governed by Poisson's equation:∇·ρ⁻¹ ∇V=I,

Where V is the voltage, ρ is the resistivity distribution and I is theimpressed current source distributions within the region being studiedand the boundary conditions are V₀ and J₀. In the case of EIT, highfrequency, low amplitude signals, e.g., 60 KHz and ˜1 microampererespectively, are used. Since there are no current sources of thisfrequency in the body, then ρ=0, and Poisson's equation becomesLaplace's equation:∇·ρ⁻¹ ∇V=0

In the field of EIT, several types of problems are studied:

-   -   1. The “forward problem”, where ρ, V₀ and J₀ are given and the        goal is to determine the voltage and current distributions V and        J.    -   2. The “inverse problem”, where V and J are given and the goal        is to determine ρ.    -   3. The “boundary value” problem where V₀ and J₀ are given and        the goal is to determine ρ, V and J.

In a preferred implementation, ρ, V and J are determined using boundaryvalue problem methods, then once ρ is determined, the optimal V₀ and J₀are determined using a modified inverse problem where the desired V andJ in and near the myocardium are given and the defibrillation waveformsfor each of the electrode pairs is generated.

In general principle, the process of EIT involves injecting a current byan electrode, and the induced voltage is measured at multiple points onthe body surface. In the preferred implementation, what is termed the“multireference method” is used for configuring the current voltagepairs. (Hua P, Webster J G, Tompkins W J 1987 Effect of the MeasurementMethod on Noise Handling and Image Quality of EIT Imaging, Proc. Annu.Int. Conf. IEEE Engineering in Medicine and Biology Society 91429-1430.) In the multireference method, one electrode is used as thereference electrode while the remaining electrodes are current sourceswith the induced voltages being measured on each electrodesimultaneously while the current is being delivered. The amplitude ofthe current sources are individually varied and each electrode istreated as a reference lead in succession. Finite element methods arethen used to convert the calculus problem (∇·ρ⁻¹∇V=0) into a linearalgebra problem of the form YV=C, where Y, V, and C are the conductance,voltage, and current matrices respectively. Y, V, and C are alsosometimes known as the master matrix, node voltage vector, and nodecurrent vector respectively. Mesh generation is performed on the two orthree-dimensional physical model with triangular or quadrilateralelements for two dimensional problems and hexahedral shapes forthree-dimensional problems. Boundary conditions are then set such as atthe reference node or driving electrodes for Dirichlet (known surfacevoltages) or Neuman (known surface currents) boundary conditions. Anumber of methods have been used to compute the master matrix such asGaussian elimination or Cholesky factorization.

The Newton-Raphson algorithm may also be used for reconstruction of theresistivity distribution. The algorithm is an iterative algorithmparticularly well suited to non-linear problems. The Newton-Raphsonmethod minimizes an error termed the “objective function”. Here, it isdefined as the equally weighted mean square difference between themeasured and estimated voltage responses:Φ(ρ)=(½)(V _(e)(ρ)−V ₀)^(T)(V _(e)(ρ)−V ₀).

Using methods known to those skilled in the art, an algorithm isutilized whereby a distribution is first estimated, then the theoreticalvoltage response to a given current input is calculated using the finiteelement method. The estimated voltages are subtracted from the measuredvoltages to obtain the objective function. If the objective function isless than an error threshold, the estimated distribution is deemed to bean acceptable estimation. If not, the following equation is used toupdate the resistivity distribution:Δρ^(k)=−[V _(e)(ρ^(k))^(T) V _(e)′(ρ^(k))]⁻¹ {V _(e)′(ρ^(k))^(T)[V_(e)′(ρ^(k))−V ₀]}

This sequence is repeated until an acceptable estimation is achieved.

In a preferred implementation, a table lookup method is provided todetermine the estimated voltage matrix V_(e)(ρ). The table values arebased on average patient resistivity distributions and assuming correctplacement of the electrode. Better accuracy can be achieved by providinganatomical markings 126 on the electrode pad as shown in FIG. 21.

Accuracy may also be improved by providing a secondary imaging methodsuch as ultrasound to take advantage of its higher imaging resolution tocalculate the positions of the internal organs relative to theelectrodes. If a secondary imaging method such as ultrasound is used todetermine the positions of internal tissues, EIT can be used todetermine the resistivities of each tissue type.

In other implementations, an average resistivity value is determined forthe tissue regions as defined by the secondary imaging method. This isaccomplished by first defining a tissue region such as the lungs ormyocardium by standard image processing methods. Next, the calculatedresistivity distribution is overlayed onto the secondary image. Allnodes of the resistivity distribution that are contained within aparticular tissue region are combined together into a single resistivitymeasure for that tissue region. The method of combination may be anaveraging, median, or other statistical or image processing method.

The optimal V₀ and J₀ are determined using a modified inverse problemwhere the desired V and J in and near the myocardium are given and thedefibrillation waveforms for each of the electrode pairs is generated.

Improved current delivery (and impedance measurements) can be achievedby close-packing a large number of electrodes. Many arrangements ofelectrodes are possible. In a preferred implementation, theconfiguration of electrodes is determined with the assistance of thetheory of tessellation. A regular tiling of polygons (in twodimensions), polyhedra (three dimensions), or polytopes (n dimensions)is called a tessellation. Tessellations can be specified using aSchläfli symbol. The breaking up of self-intersecting polygons intosimple polygons is also called tessellation, or more properly, polygontessellation. There are exactly three regular tessellations composed ofregular polyhedra symmetrically tiling the plane, as shown in FIG. 10.Tessellations of the plane by two or more convex regular polygons suchthat the same polygons in the same order surround each polygon vertexare called semi-regular tessellations, or sometimes Archimedeantessellations. In the plane, there are eight such tessellations, shownin FIG. 11. There are fourteen demi-regular (or polymorph)tessellations, which are orderly compositions of the three regular andeight semi-regular tessellations. These polyhedra are shown in FIG. 12.Other demi-regular tessellations are Penrose Tilings. In threedimensions, a polyhedron that is capable of tessellating space is calleda space-filling polyhedron. Examples include the cube, rhombicdodecahedron, and truncated octahedron. There is also a 16-sidedspace-filler and a convex polyhedron known as the Schmitt-Conwaypolyhedron, which fills space only aperiodically. Space-fillingpolyhedron can be utilized to better fit the electrodes to thethree-dimensionality of the human thorax. In the preferredimplementation, the electrode tessellation pattern is a cubic orhexagonal regular tessellation.

In another implementation, using the previously-described EIT methods,it is possible to deliver arbitrarily complex spatial and temporaldistributions of current to the heart limited only by the number andsize of electrodes on the thorax. During fibrillation, either pre orpost-shock, the direction of the activation wavefront can vary and isnot predictable, as is the case with a normal sinus rhythm. In FIG. 24B,the post-shock wavefront travels in a counter-clockwise, while in FIG.25B, the rotation is clockwise. The arrows in the figure represent thedirection of movement of the wavefronts; the shading of a regionrepresents the phase of the cardiac cycle in the region (e.g., whiterepresenting 0 msec of the cycle, and darkest gray representing 110msec). The wave may approach a region of the heart where only a smallpercentage of the cells are able to depolarize, preventing effectivemuscle contraction, but allowing the wave to continue. Indeed, the wavedoes continue, and by the time it travels completely around the heart,more cells have repolarized, and the wave can continue. The wave will,in fact, continue, and since no blood is being pumped, it is usually afatal condition unless there is an intervention with a defibrillator.The heart can support multiple such incoherent waves traveling indifferent directions, so we will apply the term wavelets to thesevarious waves. As the wavelets travel around the heart, the rate atwhich they complete a ‘circuit’ is called the VF cycle rate. As thewavelet passes a particular point and depolarizes the cells, the cellswill recover, or repolarize before the wavelet returns.

It is desirable to be able to deliver electrical current to specificregions of the myocardium so as to either reduce the extent of theexcitable gap region 92 (FIG. 25D) or to terminate pre or post-shockwavelets. The extent of the excitable gap region 92 can be reduced orabolished by delivering current sufficient to depolarize the region ofthe myocardium that would have been in the excitable gap prior to thedefibrillation shock. Using the methods of EIT just described, theresistivity distribution is calculated. Then, using a biodomain model,e.g., as described in IEEE Trans Biomed Eng 46: 260-270, 1999(Entcheva), a predicted response of the myocardium to an electricalstimulation by the device can be calculated. The bidomain model may be asystem of two reaction-diffusion equations, one for the intracellularand the other for the extracellular space linked by transmembranecurrent. The model simplifies the complex, non-linear behavior of thesarcolemmal ion channels, assuming a passive membrane behavior modeledas a parallel combination of a resistor and a capacitor in order tominimize computational complexity. The model is a system of twodifferential equations with a constant resistance membrane, Rm:Ñ×(giÑΦi)=β(Vm/Rm),Ñ×(geÑΦe)=β(Vm/Rm),where Vm and Rm are the transmembrane voltage and membrane resistance,respectively, and gi and ge are the intracellular and extracellularconductivity tensors modeling the fiber architecture of the myocardiumand β is the cell surface to volume ratio. The excitable gap region 92is located based on the bidomain calculations, and is shown in FIG. 25Das a hatched region.

Using the previously described EIT methods, the heart is stimulated byone or more pulses in succession before the defibrillation shock bycurrent that is focused in the region of the myocardium that thebidomain model predicted as occupying the excitable gap (hatched regionin FIG. 25D). Delivering the pre-shock current to those regions notadequately affected by the defibrillation shock will result in lesstotal energy being required for the combined energies of thepre-defibrillation stimulation and defibrillation shock. A tessellatedelectrode pattern may be used, e.g., incorporating 16 electrodes for theanterior and posterior electrode pad assemblies.

In another implementation, an electrode configuration is providedwhereby at least some portion of the stimulating electrodes are alsoconnected to filtering and amplification of individualelectrocardiographic (ECG) monitoring channels. In a preferredimplementation, at least 16 ECG channels are available in the devicewith the input multiplexed between one electrode on the anterior pad andone electrode on the posterior pad. A sample rate for ECG analysis ispreferably 250 Hz, therefore the A/D is sampled at 500 Hz withalternating samples from the anterior and posterior electrode. EmployingEIT and solving the forward problem as discussed previously, thedistribution of activation wavefronts on the epicardium can becalculated. The path of the activation wavefront can also cancalculated, for instance in FIGS. 24B and 25B where the arrows show thepath of the activation wavefronts. Using the previously described EITmethods, the heart is stimulated by one or more pulses in successionfollowing the defibrillation shock by current that is focused in theregion of the myocardium that lies in the path of the activationwavefront (e.g., the hatched region in FIG. 25E). Depolarizing tissue infront of the activation wavefront will quench the arrhythmogenicactivation wavefront when it reaches the refractory, recently stimulatedregion.

Referring to FIG. 1, variable resistors 71, 72 are replaced by highvoltage transistor switches and provided for each of the 16 electrodes.Eight of the 16 anterior electrodes are connected to one H-bridge 10,and the remaining eight are connected to the other H-bridge 11.Alternatively, all 16 may be connected to one H-bridge. By switching thetransistor switches 71, 72 at a rate significantly higher than theH-bridge, preferably >50 times faster, the switching of each highvoltage transistor 71, 72 can be pulse-width modulated to produce anaverage current for each electrode. As long as the switching duration isless than approximately 200 microseconds, the myocardium will respond tothe average of the pulse-width modulated (PWM) currents. As a result, itis possible also to switch the H-bridge with a period of 120microseconds, with positive and negative phases of the bridge with aduration of 55 microseconds and a 10 microsecond interphase delay. Eachindividual high voltage transistor is pulse-width modulated with pulsedurations of 5-10 microseconds during both the positive and negativephases of the bridge so as to be able to produce a potentially bipolarvoltage at any of the electrodes necessary for a fully functional EITsystem.

One possible theory to explain the improvement that some implementationsof the invention may achieve in defibrillation efficacy (understanding,of course, that the invention is not limited to this theory) is asfollows: As stated previously, the theory of Virtual ElectrodePolarization (VEP) describes the phenomena by which, because of currentflow within a partially conductive medium (the myocardium) containedwithin another partially conductive medium (blood of the cardiacchambers, lungs, interstitial fluids and other organs within thethoracic cavity), myocardial polarization during defibrillation ischaracterized by the simultaneous presence of positive and negativeareas of polarization adjacent to each other. “Phase Singularity” asdefined within the context of VEP is a critical point that is surroundedby positively polarized (equivalent to “depolarized” in the conventionalelectrophysiology nomenclature), non-polarized and negatively polarized(equivalent to “hyperpolarized”) areas. These phase singularities arethe source of re-initiation of fibrillation. Post shock excitationsinitiate in the non-polarized regions between the positively andnegatively polarized areas through a process termed “break excitation.”The break excitations propagate through the shock-induced non-polarizedregions termed “excitable gaps”, and if the positively polarized regionshave recovered excitability, then a re-entrant circuit at whichfibrillation may initiate is formed. With biphasic defibrillation, thesecond phase of the shock nullifies the VEP effect by depolarizing thenegatively polarized tissue. Since less energy is needed to depolarizerepolarized tissue than further depolarize already depolarized tissue,effective biphasic defibrillation achieves nearly completedepolarization of the myocardium by reversing the negative polarizationwhile maintaining the positive polarization. There remain, however,excitable gaps even with biphasic and multiphasic waveforms, albeitreduced in scope relative to monophasic waveforms, and there stillremains the potential for significant improvement of the efficacy ofbiphasic defibrillation waveforms.

FIG. 13 shows the results of a simulation in a study by Efimov (Am JPhysiol Heart Circ Physiol 2000; 279:H1055-70). The lighter grey region90 is a region of positive polarization and the black region 91 is oneof negative polarization. The white region 92 is the excitable gapregion. FIGS. 14a and 14b depict, in schematic view, what occurs whenthe area of the electrodes is varied. As can be seen, by increasing thesize of the electrodes, the contact angle, φ 93, of the electric fieldlines is increased in the region of the excitable gap, thereby reducingthe areal extent of the excitable gap. Reduction of the areal extent ofthe excitable gap, improves the chances for a successful defibrillationand reduces defibrillation thresholds.

In other implementations, the waveforms may each be composed of asequence of pulses. The relative timing of the current vectors may bedesigned so that the pulse sequences are interposed with non-overlappingindividual pulses.

Referring to FIGS. 26 and 27, in another implementation the electrodepairs may be configured so that, at the heart, the electrical currentvector 205 for one pair is substantially aligned with (e.g., parallel tothe long axis of) the myocardial fiber orientation of the epicardium 200on the anterior portion 202 of the heart, while the current vector 206for the other pair of electrodes is substantially aligned with the fiberorientation of the epicardium 201 on the posterior portion 203 of theheart. In a three-dimensional conductive medium, the currentdistribution is also three-dimensional, with each point within thevolume best described as a 3-space vectorial value specifying thecurrent density. Thus, the term “current vector” is a simplification ofthis more general concept, and describes a volume-averaging of thecurrent density that has been normalized to provide a vector thatdescribes the average direction of current flow within a volume.Myocardium has been shown to have anisotropic conductivity, with thelong axis along the fiber orientation having the higher conductivity.This method thus enhances current flow within the myocardium (as itlowers the impedance along paths through the myocardium, causing morecurrent to flow through the myocardium). This method is also applicableto implementations containing more than two pairs of electrodes. Forinstance, four pairs of electrodes may be used with the electricalcurrent vector for each aligned with the fiber orientation of a portionof the heart. In one implementation, the four portions of the heart arethe left anterior, right anterior, right posterior, and left posterior.

Referring to FIGS. 28A and 28B, the method described in the previousparagraph may also be applied to provide a “leadless” implanteddefibrillator system wherein all the stimulating electrodes and thedevice housing of the implanted cardiac stimulation device is externalto the thoracic cage but implanted subcutaneously. Commonly termedimplanted cardioverter-defibrillators (ICDs), the devices have requiredfor at least one of the stimulating electrodes to be placed on acatheter placed within the heart, typically in the right ventricle andsuperior vena cava in order to reduce the energy required todefibrillate, thus minimizing the size of the implanted device.

Referring to FIG. 28A, two separate defibrillator housings 207, 208 areprovided along with stimulating electrodes 209, 210, 212. Thecorresponding stimulating electrodes 209, 210 for the right atrialhousing 207, and the corresponding stimulating electrode 212 for leftatrial housing 208, result in current vectors 205, 206 substantiallyaligned with the fiber orientation of the myocardinum in the posteriorand anterior portions of the heart, respectively. The right atrialstimulating electrodes 209, 210 (which electrically serve as oneelectrode) are positioned over the intercostal spaces to reducedefibrillation impedance and are mounted onto the adjacent ribs withmounting hardware of a suitable material such as stainless steel alloys(e.g., 316L), cobalt-alloy F90, or titanium, or sutured into place.Using two pairs of stimulating electrodes provides a more uniformcurrent distribution across a larger region of the heart. Because of thehuman heart's angular orientation to the rib cage, the intercostalspaces are oriented perpendicular to the perimeter of the heart inprojection. Thus, aligning a single stimulating electrode with theintercostal space may reduce the impedance and increase therapeuticcurrent levels, but as a result myocardial current distribution becomesmore non-uniform. This current density intensification is eliminated byplacing at least two commonly connected stimulating electrodes 209, 210in different intercostal spaces. The particular intercostal spaces maybe adjacent or may have intervening, non-stimulating intercostal spaces.

Referring to FIG. 29, the housings 207, 208, which are configured as“active” cans with all or some portion of the housing providingstimulation to the stimulating electrodes, are connected to conductorscontained in a common sheath 211 composed of a biocompatible materialsuch as a silicone-polyurethane copolymer. The stimulating electrode 212corresponding to the left atrial housing 208 is located just below thexyphoid process. Preferably, this electrode is oriented parallel to theperimeter of the projection of the heart onto the rib cage to providebetter current distribution. Microprocessors 213, 214 communicate viaconductors 215, 216 to provide status information and synchronization ofdefibrillation pulses during defibrillation. Preferably, one of themicroprocessors functions as a primary processor. The primary processoris responsible for functions such as self tests and communication withdevices outside the body that are better performed in a centralizedrather than distributed manner, ECG signal processing may also belocalized to the primary processor, and the secondary processor mayperform the single function of providing a defibrillation pulse at atime deemed appropriate by the primary processor, which communicatesthat time via conductors 215, 216. Preferably, one wire 215 is asingle-wire serial interface providing a variety of status and databetween the processors, while the second wire 216 is a hardware-levelinterrupt to provide microsecond-level control of the timing of thesecondary defibrillation pulse. In a related implementation shown inFIG. 28B, conductors for electrodes 209, 210 that are part of thedefibrillation circuit contained in housing 207 pass through left atrialhousing 208, and the conductor 212 passes through housing 207, in orderto provide an alternative configuration with a minimal number of wireswithin the patient.

In another implementation, an imaging method such as ultrasound ormagneto-resonant imaging (MRI) may be used to determined the exactlocation and angular orientation of the heart within the thoracic cavityprior to implantation of the device so as to obtain improved positioningto produce current vectors 205, 206 in better alignment with the fiberorientation of the epicardium. The defibrillator housings 207, 208 andstimulating electrodes 209, 210, 212 are positioned so as to providecloser alignment of the expected current vectors 205, 206 to the fiberorientations 200, 201. In the implementation where MRI is used, externaltransthoracic pacing electrodes may be applied to the patient in thepositions such as is shown in FIG. 8a or 8 b prior to the process ofobtaining the MRI. A test pacing pulse or pulses may then be deliveredto the patient that is synchronized with the MRI excitation pulses so asto produce a three dimensional representation of the current vectorspace that will show the myocardial fiber orientation for use indetermining the positioning of the housings 207, 208 and stimulatingelectrodes 209,210, 212.

In another implementation, resistance circuits 55, 56 are eliminated andthe waveform shape, and thus also the first phase average current, isadjusted by pulse width modulating the switches in the H-bridges 10, 11.This configuration is the Class D amplifier configuration, known tothose skilled in the art of amplifier design. In its simplest form, aswitch-mode amplifier consists of an H-bridge and a load as shown inFIG. 15. Amplifiers are typically classified by their output stages. Ofthe common output-stage topologies (Classes A, B, AB, and D), Class Damplifiers exhibit the highest efficiency. A linear output stage (ClassA, B, or AB) draws considerable bias current while sourcing and sinkingcurrent into a speaker making them not particularly well suited to highvoltage designs. A nonlinear (Class D) output stage eliminates this biascurrent. In the preferred implementation, as shown in FIG. 16, the ClassD amplifier consists of an input preamplifier 95 for isolating,filtering and level shifting the control voltage from the processingmeans 5, a sawtooth oscillator 96, a comparator 97, two MOSFET drivers98, 99, and the H-bridge switches 100-103. The comparator samples theinput signal, with the oscillator frequency determining the duration ofthe sampling period. Thus, the oscillator frequency is an importantfactor in the overall performance of a Class D amplifier. As shown inFIG. 17, the comparator output 104 is a pulse-width modulated (PWM)square wave that drives the H-bridge. The PWM square wave 104 is createdby a comparator whose inputs are the sawtooth (V_(RAMP)) 105 and thecontrol signal (V_(IN)) 106. The H-bridge then outputs the square wavedifferentially. For a given input level, the comparator output is aduty-cycle modulated square wave with period determined by the sawtoothfrequency. The PWM square wave controls the H-bridge drivers 100-103,turning opposite pairs of MOSFETs off and on, thereby reversing currentto the load within a single period. The output may be filtered bycapacitor filters or inductor/capacitor filter combinations which removehigh-frequency content from the H-bridge square wave output.

Alternatively, the measurement of the thoracic cavity may be carried outusing an ultrasound transducer capable of imaging the heart andsurrounding tissue. An ultrasound transducer may be incorporated into anintegrated defibrillation pad, as shown in FIG. 18. In a preferredimplementation, an opening in the center of the electrode is providedthat is covered over with an ultrasonic-conducting gel 107. The gel is abilayer structure with a more aggressive adhesive provided on the faceopposite to the patient for attaching the ultrasonic probe prior to use.

In other implementations, there may be two or more separatedefibrillators, as shown in FIG. 19. The first defibrillator 110 acts asthe master defibrillator, while additional defibrillators 111 functionas slave defibrillators whose energy is delivered synchronously withthat of the master defibrillator 110. Synchronization is provided bycommunication means 114. Preferably, the communication means 114 isimplemented as a simple switch. In a conventional defibrillator, thedelivery of energy is initiated via the closure of a discharge switch112 located on the front panel or on a set of defibrillation paddles.The closure of the switch initiates the defibrillation sequence underthe control of processing means 5. Charging of the high voltagecapacitors on both defibrillators 110, 111 is initiated via thecharge-control user inputs 115. At the appropriate time, the clinicianwill press the discharge button 112. This causes the processing means 5on the first defibrillator 110 to close a slave discharge switch thatinitiates the discharge sequence on the second defibrillator 111, atwhich time the first defibrillator 110 also initiates its dischargesequence. The wiring for the communication means 114 is preferablyconfigured such that the wires are located within the same cable as theenergy delivery wires, thus reducing any additional cabling. Thecommunication means 114 may also incorporate digital communicationmethods which provide additional information about defibrillator status.

The defibrillator pad 123 may integrate all connections into a singleconnector 120 as shown in FIG. 20. The defibrillator pads may beconstructed such that a seam line 121 is located between the activeareas 122 of the pad 123 where the seam line 121 is of higher compliancethan the active areas such that the pad 123 can be folded during storagewithout creasing the active areas.

In another implementation, a physiological parameter, e.g., theelectrocardiograph (ECG), is measured in conjunction with the EIT image,and an estimate is made by the device of the chances for a successfuldefibrillation shock based analysis of ECG data. Depending on theestimate of shock success, decisions as to the proper treatment toprovide the patient are made in a coordinated resuscitation effort thatincludes both defibrillation and chest compressions, which can beprovided manually in response to prompts, or in a semi-automated orfully automated fashion. The block diagram and flow chart for such asystem is shown in FIGS. 22 and 23.

One or more additional electrodes 125 may be provided for diaphragmaticstimulation (DS) and may be incorporated into the anterior electrodesuch that the DS Electrode (DSE) is located over the patient's diaphragmas shown in FIG. 21. Diaphragmatic stimulation induces air exchange inthe lungs during cardiopulmonary resuscitation (CPR) for improvedoxygenation. The return path for the stimulation current from the DSE isthrough one of the pre-existing electrodes. Utilizing EIT or otherimaging methods, the current distribution may be adjusted to achieveoptimal stimulation, as described previously in this patent. The DSE maybe integrated with defibrillation and cardiac pacing to provide acoordinated resuscitation effort in an automated or semi-automatedfashion. The integrated resuscitation may also incorporate a means ofproviding chest compressions, such as a piston-based system manufacturedby Michigan Instruments (Michigan) or a constricting band systemmanufactured by Revivant Corp. (California). FIG. 23 shows a decisionflow chart of one possible integrated resuscitation protocol.

Many other implementations of the invention other than those describedabove are within the invention, which is defined by the followingclaims. The invention applies to both defibrillation and cardioversion;in the claims, references to defibrillation should be interpreted asalso encompassing cardioversion. Some implementations of the inventionare broader than defibrillation and cardioversion.

What is claimed is:
 1. A defibrillation system for synchronizedcardioversion of a patient comprising: a first housing comprising: afirst shock delivery circuit configured to be controlled by a firstprocessor for delivery of a first shock; a measurement circuitconfigured to receive electrocardiogram (ECG) signals and measure ECGparameters based on the ECG signals, and the first processor configuredto: analyze the ECG parameters, and initiate communication of asynchronization signal for a second processor for delivery of a secondshock; and a second housing that is separate from and external to thefirst housing, the second housing comprising: a second shock deliverycircuit, and the second processor, wherein the second processor isconfigured to: receive the communication of the synchronization signalfrom the first processor, and control the second shock delivery circuitto deliver the second shock in response to the synchronization signal.2. The defibrillation system of claim 1 wherein the delivery of thesecond shock is based on the ECG parameters.
 3. The defibrillationsystem of claim 1 comprising: three or more electrodes configured toattach to the patient and provide two or more discharge pathways; and atleast one measurement circuit configured to: couple to the three or moreelectrodes, and measure one or more electrical parameters of thepatient, wherein at least one of the first processor and the secondprocessor is configured to: based on the one or more measured electricalparameters, determine internal impedances, for the patient,corresponding to the two or more discharge pathways, and control atleast one parameter of at least one of the first shock and the secondshock based on the determined internal impedances.
 4. The defibrillationsystem of claim 3 wherein the one or more electrical parameters comprisea current that passes through the patient.
 5. The defibrillation systemof claim 3 wherein the one or more electrical parameters comprise aninduced voltage at a plurality of locations on the patient.
 6. Thedefibrillation system of claim 3 wherein the at least one of the firstprocessor and the second processor is configured to independentlycontrol the at least one parameter for each of the first shock and thesecond shock based on the internal impedances.
 7. The defibrillationsystem of claim 3 wherein the at least one parameter comprises at leastone of amplitude, tilt, duration first phase duration, second phaseduration, current, voltage, and first phase average current.
 8. Thedefibrillation system of claim 3 wherein at least the first processor isconfigured to use electrical impedance tomography (EIT) and determinethe internal impedances as an internal impedance distribution.
 9. Thedefibrillation system of claim 8 wherein one or more of the firstprocessor and the second processor is configured to determine apredicted response of the myocardium of the patient to at least one ofthe first shock and the second shock based on a model and the internalimpedance distribution.
 10. The defibrillation system of claim 9 whereinthe model comprises a biodomain model.
 11. The defibrillation system ofclaim 9 wherein the predicted response comprises an indication of one ormore excitable gap regions of the myocardium of the patient.
 12. Thedefibrillation system of claim 11 wherein at least the second processoris configured to control at least one of the first shock deliverycircuit and the second shock delivery circuit to deliver one or moreelectrical pulses in succession prior to at least one of the first shockand the second shock wherein a current density from the one or moreelectrical pulses is configured to be directed towards the one or moreexcitable gap regions of the myocardium of the patient via apredetermined physical arrangement of the three or more electrodes. 13.The defibrillation system of claim 3 wherein the measurement circuitcomprises individual ECG monitoring channels and further wherein atleast a portion of the three or more electrodes are coupled to theindividual ECG monitoring channels.
 14. The defibrillation system ofclaim 13 wherein one or more of the first processor and the secondprocessor is configured to determine an epicardial activation wavefrontdistribution.
 15. The defibrillation system of claim 13 wherein one ormore of the first processor and the second processor is configured todetermine an epicardial activation wavefront path.
 16. Thedefibrillation system of claim 13 comprising at least sixteen ECGmonitoring channels and at least sixteen electrodes.
 17. Thedefibrillation system of claim 16 wherein the first processor isconfigured to sample ECG signals at a sample rate of approximately 250Hz.
 18. The defibrillation system of claim 13 wherein the three or moreelectrodes are distributed between at least one anterior electrode padassembly and at least one posterior electrode pad assembly and furtherwherein the one or more of the first processor and the second processoris configured to sample the ECG signals with alternating samples fromthe at least one anterior electrode pad assembly and the at least oneposterior electrode pad assembly.